1. Field of the Invention.
The field of the invention is magnetic resonance imaging (MRI) and, in particular, local coils for use in receiving MRI signals.
2. Background Art.
A. MRI Imaging
In MRI, a uniform magnetic field B.sub.0 is applied to an imaged object along the z axis of a Cartesian coordinate system, the origin of which is approximately centered within the imaged object. The effect of the magnetic field B.sub.0 is to align the object's nuclear spins along the z axis.
In response to a radio frequency (RF) excitation signal of the proper frequency, oriented within the x-y plane, the nuclei precess about the z-axis at their Larmor frequencies according to the following equation: EQU .omega.=.gamma.B.sub.0
where .omega. is the Larmor frequency, and .gamma. is the gyromagnetic ratio which is constant and a property of the particular nuclei.
Water, because of its relative abundance in biological tissue and the properties of its nuclei, is of principle concern in such imaging. The value of the gyromagnetic ratio .gamma. for water is 4.26 kHz/gauss and therefore, in a 1.5 Tesla polarizing magnetic field B.sub.0, the resonant or Larmor frequency of water is approximately 63.9 MHz.
In a typical imaging sequence for an axial slice, the RF excitation signal is centered at the Larmor frequency .omega. and applied to the imaged object at the same time as a magnetic field gradient G.sub.z is applied. The gradient field G.sub.z causes only the nuclei, in a slice through the object along a x-y plane, to have the resonant frequency .omega. and to be excited into resonance.
After the excitation of the nuclei in this slice, magnetic field gradients are applied along the x and y axes. The gradient along the x axis, G.sub.x, causes the nuclei to precess at different frequencies depending on their position along the x axis, that is, G.sub.x spatially encodes the precessing nuclei by frequency. The y axis gradient, G.sub.y, is incremented through a series of values and encodes the y position into the rate of change of phase of the precessing nuclei as a function of gradient amplitude, a process typically referred to as phase encoding.
A weak nuclear magnetic resonance generated by the precessing nuclei may be sensed by the RF coil and recorded as an NMR signal. From this NMR signal, a slice image may be derived according to well known reconstruction techniques. An overview NMR image reconstruction is contained in the book "Magnetic Resonance Imaging, Principles and Applications" by D. N. Kean and M. A. Smith.
B. Local Coils
The quality of the image produced by MRI techniques is dependent, in part, on the strength of the NMR signal received from the precessing nuclei. For this reason, it is known to use an independent RF receiving coil placed in close proximity to the region of interest of the imaged object to improve the strength of this received signal. Such coils are termed "local coils" or "surface coils". The smaller area of the local coil permits it to accurately focus on NMR signals from the region of interest. Further, the RF energy of the field of such a local coil is concentrated in a smaller volume giving rise to improved signal-to-noise ratio in the acquired NMR signal.
The signal-to-noise ratio of the NMR signal may be further increased by orienting two coils at 90.degree. angles about the imaged object so that each detects RF energy along one of a pair of mutually perpendicular axes. This technique is generally known as quadrature detection and the signals collected are termed quadrature signals.
The outputs of the quadrature coil pairs are combined so as to increase the strength of the received signal according to the simple sum of the output signals from the coils. The strength of the uncorrelated noise component of these signals, however, will increase only according to the square root of the sum of the noise voltage increases times the .sqroot.2 of the noise components. As a result, the net voltage signal-to-noise ratio of the combined quadrature signals increases by approximately .sqroot.2 over the voltage signal-to-noise ratio of the individual signals.
The quadrature orientation of the two coils introduces a 90.degree. phase difference between the NMR signals detected by these coils. Therefore, combining the outputs from the two quadrature coils, to achieve the above described signal-to-noise ratio improvement, requires that one signal be shifted to have the same phase as the other signal so that the amplitudes of the signals simply add. Such phase shifting and combining is typically accomplished by means of a hybrid network. Hybrid networks are four-port networks known in the art and having the property that when the four ports are properly terminated, energy input to two of the ports, with the proper relative phase angles, will be combined at one of the remaining two ports. The antenna coils are attached to two of the ports and the output lead is attached to a third port. The remaining uncommitted port is connected to a termination resistor.
As used herein, the term quadrature coil and quadrature signal, will refer to the detecting of the NMR signal along multiple axes and combining the signals so collected, with the appropriate phase shifts to produce a signal of improved signal-to-noise ratio.
C. Planar Coils
The use of quadrature coils of conventional design may be undesirably constraining to the patient who must be positioned between two coils within the relatively small volume of the magnet bore. Further, in order that the local coil may be conveniently located on the patient, it is necessary that the quadrature local coil be substantially open on one side.
It is known, therefore, for certain imaging applications such as the imaging of the spine, to construct a local coil on a substantially planar cradle to be attached to the upper surface of the patient support table so that the patient may simply lie on top of the coil and so that the coils structure is not unduly constraining. Such open coils are termed "planar" coils to distinguish them from "whole volume" coils such as might be constructed of opposed saddle coils or solenoids. The prior art has recognized the desirability of a quadrature, planar coil. See, for example, U.S. Pat. No. 4,721,913 issued January 26 to Hyde et al.
As shown in FIGS. 1 and 2, such quadrature, planar coils 10 may employ two components, a butterfly coil 12 having two loops 14 and 16 to be sensitive to horizontal components of flux 18 within a region of sensitivity 20, and an ordinary loop coil 22 centered on the butterfly to detect vertical components of flux 24 within region 20. The signal from the two loops 14 and 16 of the butterfly coil 12 is combined with the signal from the loop coil 22, after one is shifted by 90.degree., to form a quadrature signal 17.
The two loops 14 and 16 of the butterfly coil 12 are sensitive only to countercyclic current flows as a result of the geometry of their connection. Current flow is indicated by the arrows on loops 14 and 16 in FIG. 1 and the cross superimposed on the cross-section of the loops 14 and 16 in FIG. 3, which, per convention indicates a receding flow of current. Accordingly, the butterfly coil 12 is generally insensitive to fluctuating uniform magnetic fields which induce cancelling co-cyclic currents in each of the loops 14 and 16.
An improvement to this design employing a butterfly coil 12 and loop coil 22 is taught in U.S. Pat. No. 5,030,915, issued Jul. 9, 1991 to Boskamp, the same inventor as that of the present invention, this patent incorporated herein by reference. In the Boskamp design, the butterfly coil 12 and loop coil 22 is realized by superimposed currents on a single coil structure, that structure being generally a loop bisected by a single conductor. This single structure provides improved resistance to misalignment.